CN114901321A

CN114901321A - Ceramic support

Info

Publication number: CN114901321A
Application number: CN202080091552.1A
Authority: CN
Inventors: 李向佳; 陈勇; 柴洋
Original assignee: Alfred E Mann Institute for Biomedical Engineering of USC
Current assignee: University of Southern California USC
Priority date: 2019-11-01
Filing date: 2020-10-30
Publication date: 2022-08-12
Also published as: US20240157024A1; EP4051326A1; WO2021087335A1

Abstract

The present disclosure generally relates to ceramic stents. The present disclosure relates particularly to ceramic scaffolds for bone regeneration. The present disclosure also relates to ceramic scaffolds comprising Hydroxyapatite (HA), tricalcium phosphate (TCP), or mixtures thereof. The present disclosure also relates to ceramic scaffolds having high mechanical strength and flexibility. The present disclosure further relates to ceramic scaffolds fabricated by a three-dimensional (3D) printing method, methods of fabricating ceramic scaffolds, and methods of replacing a bone of a subject using the ceramic scaffolds.

Description

Ceramic support

Cross Reference to Related Applications

This application claims the benefit of U.S. provisional patent application 62/929,630 entitled "Ceramic Scaffold" filed on 11/1/2019, attorney docket number AMISC.012PR, which is incorporated herein by reference in its entirety.

Technical Field

The present disclosure generally relates to ceramic stents. The present disclosure relates particularly to ceramic scaffolds for bone regeneration. The present disclosure also relates to ceramic scaffolds comprising Hydroxyapatite (HA), tricalcium phosphate (TCP), or mixtures thereof. The present disclosure also relates to ceramic scaffolds having high mechanical strength and flexibility. The present disclosure further relates to ceramic scaffolds fabricated by three-dimensional (3D) printing methods.

Background

Over millions of patients worldwide per year require surgical reconstruction of bone defects caused by trauma, tumors or infections [1-3 ]. Common therapeutic approaches (including autologous bone grafts [4], allogeneic bone grafts [4] and replacement material-based bone grafts [5]) still have certain limitations in treating bone injuries [6 ]. For example, autograft can only be applied to small areas of bone defects due to quantity limitations, and it can also lead to a high risk of acquiring morbidity [7 ]. With the development of tissue engineering, a new bone defect treatment method (the construction of bone tissue microenvironment by integration of extracellular matrix, cells, growth factors, etc.) has been proposed for regenerating new bone for the healing of bone defects [8 ]. To reproduce natural living matter, 3D scaffolds were further designed and manufactured using biocompatible and biodegradable materials [9 ]. 3D scaffolds with complex internal microstructures [10,11] can be achieved with the benefit of improved Additive Manufacturing (AM) manufacturing capabilities. With the advancement of AM technology, more and more materials, such as bioceramics, polymers, hydrogels, and nanocomposites, can be replicated to mimic natural bone tissue [12 ]. Of these biomaterials, bioceramic-like Hydroxyapatite (HA) and tricalcium phosphate (TCP) exhibit promising properties in cell attachment and proliferation [13], osteoconduction [14], osteointegration [15] and osteoinduction [16 ]. However, HA/TCP scaffolds made by conventional methods [8-9] are too brittle to be further manipulated. Poor mechanical properties limit the range of applications of bioceramic scaffolds in bone tissue regeneration. It is known that the compressive strength of natural trabecular bone and cortical bone reaches 100-130MPa and 130-190MPa respectively [17], and some studies indicate that 3D scaffolds with similar bone mechanical properties can promote bone regeneration [17-20 ]. Various studies have been carried out to improve the mechanical properties of HA/TCP scaffolds, for example, by increasing the sintering temperature [21], or by combining with other components [22 ]. However, while mechanical strength is increased, degradation of the scaffold is also slowed, which may lead to more serious infections and elimination risks [23 ]. Ideally, the 3D printed scaffold should gradually degrade during new bone regrowth, and at the same time the remaining 3D scaffold can still maintain a shape with some mechanical strength [24, 25 ]. In general, most stents manufactured using current manufacturing methods do not meet the trade-off between mechanical properties and degradation.

Disclosure of Invention

In the present disclosure, a ceramic stent may include a frame and a coating. The framework may comprise Hydroxyapatite (HA), tricalcium phosphate (TCP), or a mixture thereof. The coating may comprise a polymer ("coating polymer").

In the present disclosure, the frame may have at least one surface. The coating may be formed on at least one surface of the frame. The coating may at least partially cover at least one surface of the frame. Alternatively, the coating may substantially cover at least one surface of the frame.

In the present disclosure, the coating polymer may include a polymer, which may be formed by using surgical glue, a gel-like protein mixture, poly (ethylene glycol) dimethacrylate (PEGDMA), methacrylated gelatin (GelMA), gelatin, or a mixture thereof. Alternatively, the coating polymer may include a polymer formed by using surgical glue. In some embodiments, the surgical glue polymer is 2-octyl cyanoacrylate (Dermabond). Alternatively, the coating polymer may comprise gelatin. Alternatively, the coating polymer may include a polymer formed by using cyanoacrylate glue, fibrin sealant, collagen-based compound, glutaraldehyde glue, hydrogel, or a mixture thereof. Alternatively, the coating polymer may include a polymer formed by using n-butyl cyanoacrylate, 2-octyl cyanoacrylate, or a mixture thereof.

In the present disclosure, the mechanical strength of the ceramic scaffold may be in the range of 15N to 33N when measured as the maximum load of the stress-strain curve. For example, the mechanical strength may be 15N, 16N, 17N, 18N, 19N, 20N, 21N, 22N, 23N, 24N, 25N, 26N, 27N, 28N, 29N, 30N, 31N, 32N, or 33N or any range defined by these values. The ceramic scaffold may have a flexural strength in the range of 10MPa to 50MPa, including 10MPa, 11MPa, 12MPa, 13MPa, 14MPa, 15MPa, 16MPa, 17MPa, 18MPa, 19MPa, 20MPa, 21MPa, 22MPa, 23MPa, 24MPa, 25MPa, 26MPa, 27MPa, 28MPa, 29MPa, 30MPa, 31MPa, 32MPa, 33MPa, 34MPa, 35MPa, 36MPa, 37MPa, 38MPa, 39MPa, 40MPa, 41MPa, 42MPa, 43MPa, 44MPa, 45MPa, 46MPa, 47MPa, 48MPa, 49MPa, or 50MPa or any range defined by these values; where the bending strength is measured by a standard three-point bending test. Alternatively, the mechanical strength of the ceramic support may be at least 5 times higher than the mechanical strength of the frame. Alternatively, the mechanical strength of the ceramic support may be at least 10 times greater than the mechanical strength of the frame. Alternatively, the mechanical strength of the ceramic support may be 10 to 20 times higher than that of the frame.

In the present disclosure, the thickness of the coating polymer may be in the range of 1 micron to 1000 microns. Alternatively, the thickness of the coating polymer may be in the range of 10 microns to 500 microns.

The present disclosure also relates to methods of making ceramic stents. The method can comprise the following steps: preparing a slurry comprising Hydroxyapatite (HA), tricalcium phosphate (TCP), or a mixture thereof; and a UV polymerizable monomer formulation; preparing a green body using a three-dimensional (3D) printing method and a slurry; debinding the green body and sintering to remove polymer formed by polymerization of the UV polymerizable monomer formulation to produce a sintered porous body, wherein the sintered porous body forms a framework; coating the sintered porous body with a polymer coating solution; polymerizing the polymer coating solution to form a coating comprising a polymer ("coating polymer"); and thereby obtaining a ceramic support.

In the present disclosure, the three-dimensional (3D) printing method may be a paste printing based on mask image projection (MIP-SP) method.

In the present disclosure, the ceramic scaffold may be sintered at a temperature in the range of 1000 degrees celsius to 1500 degrees celsius. Or wherein the green body may be sintered at a temperature in a range of 1050 degrees celsius to 1250 degrees celsius.

In the present disclosure, the coating of the sintered porous body may be performed by a method that may include a surface spray method, a brush spray method, a vacuum fusion method, or a combination thereof.

Any combination of the above ceramic scaffolds and/or methods of making these ceramic scaffolds is within the scope of the present disclosure.

Some aspects relate to a ceramic stent, comprising:

a framework comprising Hydroxyapatite (HA), tricalcium phosphate (TCP), or a mixture thereof; and

a coating comprising a coating polymer;

wherein the frame has at least one surface, wherein a coating is formed on the at least one surface of the frame, and wherein the coating at least partially covers the at least one surface of the frame.

In some examples, the coating polymer includes a polymer formed by using surgical glue, a colloidal protein mixture, poly (ethylene glycol) dimethacrylate (PEGDMA), methacrylated gelatin (GelMA), gelatin, or a mixture thereof.

In some examples, the coating polymer comprises surgical glue, gelatin, or a mixture thereof.

In some examples, the coating polymer includes a polymer of n-butyl cyanoacrylate monomer, 2-octyl cyanoacrylate monomer, or a mixture thereof.

In some examples, the coating is coated on the surface of the stent at a thickness of 5 μm to 1mm

In some examples, the ceramic scaffold has a mechanical strength in the range of 15N to 33N when measured as the maximum load of a stress-strain curve.

In some examples, the ceramic scaffold has a flexural strength in the range of 10MPa to 50 MPa; where the bending strength is measured by a standard three-point bending test.

In some examples, the mechanical strength of the ceramic scaffold is 10 to 20 times greater than the mechanical strength of the frame.

In some examples, the coating polymer includes a polymer formed by using cyanoacrylate glue, fibrin sealant, collagen-based compounds, glutaraldehyde glue, hydrogel, or a mixture thereof.

In some examples, the coating polymer includes an acrylate polymer.

In some examples, the coating polymer includes a cyanoacrylate polymer.

In some examples, the coating polymer comprises a surgical glue.

In some examples, the coating polymer comprises gelatin.

In some examples, the scaffold comprises stacked layers of Hydroxyapatite (HA), tricalcium phosphate (TCP), or a mixture thereof, each layer having a thickness of 10 μm to 200 μm.

In some examples, the mechanical strength of the ceramic scaffold is at least 5 times greater than the mechanical strength of the frame.

In some examples, the mechanical strength of the ceramic scaffold is at least 10 times greater than the mechanical strength of the frame.

In some examples, the coating polymer has a thickness in a range from 1 micron to 1000 microns.

In some examples, the coating polymer has a thickness in a range of 10 microns to 500 microns.

In some examples, the coating polymer includes a polymer formed by using surgical glue, a colloidal protein mixture, poly (ethylene glycol) dimethacrylate (PEGDMA), methacrylated gelatin (GelMA), gelatin, or a mixture thereof; wherein the ceramic support has a mechanical strength in the range of 15N to 33N and/or a bending strength in the range of 10MPa to 50MPa when measured as the maximum load of the stress-strain curve; and wherein the bending strength is measured by a standard three-point bending test.

In some examples, the coating polymer includes a polymer formed by using surgical glue, a gel-like protein mixture, poly (ethylene glycol) dimethacrylate (PEGDMA), methacrylated gelatin (GelMA), gelatin, or a mixture thereof; wherein the mechanical strength of the ceramic scaffold is at least 5 times higher than the mechanical strength of the frame, or at least 10 times higher than the mechanical strength of the frame, or 10 to 20 times higher than the mechanical strength of the frame; and wherein the mechanical strength of the ceramic scaffold is the maximum load of the stress-strain curve.

In some examples, the coating polymer includes a polymer formed by using n-butyl cyanoacrylate, 2-octyl cyanoacrylate, or a mixture thereof; wherein the ceramic support has a mechanical strength in the range of 15N to 33N and/or a bending strength in the range of 10MPa to 50MPa when measured as the maximum load of the stress-strain curve; and wherein the bending strength is measured by a standard three-point bending test.

In some examples, the coating polymer includes a polymer formed by using n-butyl cyanoacrylate, 2-octyl cyanoacrylate, or a mixture thereof; wherein the mechanical strength of the ceramic scaffold is at least 5 times higher than the mechanical strength of the frame, or at least 10 times higher than the mechanical strength of the frame, or 10 to 20 times higher than the mechanical strength of the frame; and wherein the mechanical strength of the ceramic scaffold is the maximum load of the stress-strain curve.

In some examples, the coating substantially covers at least one surface of the frame.

In some examples, the coating polymer includes a polymer formed by using a surgical glue.

In some examples, the coating polymer includes a polymer formed by using gelatin.

In some examples, the coating polymer includes a polymer formed by using cyanoacrylate glue, fibrin sealant, collagen-based compound, glutaraldehyde glue, hydrogel, or a mixture thereof.

In some examples, the coating polymer includes a polymer formed by using n-butyl cyanoacrylate, 2-octyl cyanoacrylate, or a mixture thereof.

In some examples, the coating polymer includes a polymer formed by using surgical glue, a colloidal protein mixture, poly (ethylene glycol) dimethacrylate (PEGDMA), methacrylated gelatin (GelMA), gelatin, or a mixture thereof; wherein the mechanical strength of the ceramic scaffold is at least 5 times higher than the mechanical strength of the frame, or at least 10 times higher than the mechanical strength of the frame, or 10 to 20 times higher than the mechanical strength of the frame; and wherein the mechanical strength of the ceramic scaffold is the maximum load of the stress-strain curve.

Some aspects relate to a method of manufacturing a ceramic stent, comprising:

preparing a slurry comprising Hydroxyapatite (HA), tricalcium phosphate (TCP), or a mixture thereof; and a UV polymerizable monomer formulation;

preparing a green body using a three-dimensional (3D) printing method and a slurry;

debinding the green body and sintering to remove polymer formed by polymerization of the UV polymerizable monomer formulation to produce a sintered porous body, wherein the sintered porous body forms a framework;

coating the sintered porous body with a polymer coating solution;

polymerizing the polymer coating solution to form a coating comprising a coating polymer; and thereby

Obtaining the ceramic bracket.

In some examples, the three-dimensional (3D) printing method is a paste printing based on mask image projection (MIP-SP) method.

In some examples, the green body is sintered at a temperature in a range of 1000 degrees celsius to 1500 degrees celsius.

In some examples, the green body is sintered at a temperature in a range of 1050 degrees celsius to 1250 degrees celsius.

In some examples, the coating of the sintered porous body is performed by a method including a surface spray coating method, a brush spray coating method, a vacuum fusion method, or a combination thereof.

In some examples, the obtained ceramic scaffold is the ceramic scaffold of any one of claims 1 to 45.

In some examples, the coating of the sintered porous body is performed by a method including a surface spray method, a brush spray method, a vacuum fusion method, or a combination thereof.

In some examples, degreasing is performed under vacuum at a pressure of 0.01-0.5 MPa.

In some examples, the coating is applied to the stent under vacuum.

In some examples, the coating is applied to the stent by surface spraying.

Some aspects relate to a method of replacing a bone of a subject, comprising:

identifying a portion of bone missing from the subject or removing a portion of bone from the subject, an

Placing the ceramic scaffold of any of claims 1-8 in place of the missing or removed bone portion, wherein the ceramic scaffold allows bone regeneration within and around the ceramic scaffold.

Some aspects relate to a method of replacing a bone of a subject, comprising:

Placing the ceramic scaffold of any one of claims 1-45 in a location of the missing or removed bone portion, wherein the ceramic scaffold allows bone regeneration within and around the ceramic scaffold.

Any combination of the above aspects and examples is within the scope of our claims.

These and other components, steps, features, objects, benefits and advantages will become apparent from a review of the following detailed description of illustrative examples, the accompanying drawings and the claims.

Drawings

The drawings are illustrative examples. They do not illustrate all examples. In addition or alternatively, other examples may be used. Details that may be obvious or unnecessary may be omitted to save space or for a more efficient illustration. Some examples may be implemented with other components or steps and/or without all of the components or steps shown. When the same number appears in different drawings, it refers to the same or like components or steps.

The patent or application file contains at least one drawing executed in color. Copies of this patent or patent application publication with color drawing(s) will be provided by the office upon request and payment of the necessary fee.

FIG. 1 manufacturing process of enhanced HA/TCP scaffold for bone defects. (a) Fractures and severe defects; (b) the components of HA/TCP slurry; (c) schematic diagram of MIP-SP method of making HA/TCP green body; (d)3D printing a green body of the HA/TCP support; (e) a brown blank of the HA/TCP bracket after the degreasing process; (f) pure HA/TCP stent after sintering process; (g) schematic diagram of vacuum coating process; (h) HA/TCP stents after the coating process; and (i) tailoring the HA/TCP scaffold to the shape of the fracture and severe defect.

FIG. 2 effect of coating material on mechanical properties of HA/TCP microcell structure. (a) Schematic representation of a second stage coating process by using vacuum fusion; (b, c) full view and Scanning Electron Microscope (SEM) images of HA/TCP micro-units without coating and with different coating materials; and (d) maximum load comparison of HA/TCP microcell structures with and without coating.

FIG. 3 mechanically reinforces HA/TCP stent by post-coating method. (a) Schematic diagram of post-coating process; (b-c) full view and SEM (scanning electron microscope) images of HA/TCP stent before and after the coating process; (d) sectional views of HA/TCP microcell structures with different coating methods and coating parameters; (e) the maximum load that the HA/TCP microcell structure can bear after being enhanced by different coating methods; (f) fracture toughness of HA/TCP printed parts treated with different coating methods; and (g) shape change of the 3D printed HA/TCP part after the sintering and coating process.

FIG. 4 compression characteristics of3D printed HA/TCP components before and after the coating process. (a) After the coating process, the coating material is filled in the pores between the HA/TCP particles; (b-c) simulation results of stress distribution of3D printed HA/TCP parts before and after the coating process using COMSOL Multiphysics, respectively; (d) the cracks are only deflected along the HA/TCP particles without coating material; (e) simulation results of stress distribution of3D printed HA/TCP components after the coating process sintered at different temperatures, respectively, by using COMSOL Multiphysics; (f) compressive properties of3D printed HA/TCP microcell structures sintered at 1250 ℃ with and without a surgical glue coating; and (g) compression characteristics of3D printed HA/TCP parts sintered at different temperatures with and without a surgical glue coating.

FIG. 5 comparison of flexural strength, fracture toughness, and tensile strength of3D printed HA/TCP components before and after the coating process. (a) Three-point bending test of HA/TCP printing parts; (b-D) load displacement, fracture toughness and flexural strength of3D printed HA/TCP parts with and without coating material; (e) simulation of stress distribution of3D printed HA/TCP parts with and without coating material by COMSOL Multiphysics; SEM images showing fracture surfaces of pure 3D printed HA/TCP part (f) and the surgical glue coated 3D printed HA/TCP part (g); (h)3D printing a tensile test of the HA/TCP component; (i, j) strain-stress and young's modulus of3D printed HA/TCP parts sintered at different temperatures with and without coating material; and (k) simulated appearance of3D printed HA/TCP parts sintered at low and high temperatures, respectively, by using COMSOL Multiphysics.

FIG. 63D tailors the design and manufacture of HA/TCP boards. (a)3D can trim the CAD model of the HA/TCP board; 3D printing images of HA/TCP boards before (b) and after (c) degreasing and sintering; (d) having a surgical glue coating; (e) a finishing method of HA/TCP board; (f) HA/TCP plates for the cutting of craniofacial defects; (g) SEM image of HA/TCP plate; and (h) mounting the metal screws on the trimmable HA/TCP board of the jig.

FIG. 7 reconstruction of craniofacial bones with a coating enhanced 3D printed HA/TCP scaffold. (a) Craniofacial bone defect animal models; (b) a CAD model of HA/TCP scaffolds for severe craniofacial bone defects; degreasing and sintering images of the 3D printed HA/TCP scaffold before (c) and after (D); (e) having a surgical glue coating; (f) compression simulation of HA/TCP scaffolds using COMSOL Multiphysics; and (g) compression testing of3D printed HA/TCP scaffolds without a coating and (h) with a coating; and (i) force and displacement in compression testing of HA/TCP scaffolds with and without coatings.

FIG. 8 design and fabrication of HA/TCP scaffolds for long bone defect reconstruction. (a) A digital model of a mouse severe femoral defect animal model; (b) a CAD model of a scaffold for a long bone defect; 3D printing images of the HA/TCP scaffold before (c) and after (D) degreasing and sintering; (e) compression simulation of HA/TCP scaffolds using COMSOL Multiphysics; (f) 3D printing images of HA/TCP stent with surgical glue coating; and (g) a comparison of the maximum loads that can be withstood by HA/TCP stents with different manufacturing parameters.

Detailed Description

Illustrative examples are now described. In addition or alternatively, other examples may be used. Details that may be obvious or unnecessary may be omitted to save space or for a more efficient description. Some examples may be implemented with other components or steps and/or without all of the components or steps shown.

The present disclosure relates generally to ceramic stents. The present disclosure relates particularly to ceramic scaffolds for bone regeneration. The present disclosure also relates to ceramic scaffolds comprising Hydroxyapatite (HA), tricalcium phosphate (TCP), or mixtures thereof. The present disclosure also relates to ceramic scaffolds having high mechanical strength and flexibility. The present disclosure further relates to ceramic scaffolds fabricated by a three-dimensional (3D) printing method.

In the present disclosure, the coating polymer may include a polymer that may be formed by using surgical glue, a colloidal protein mixture, poly (ethylene glycol) dimethacrylate (PEGDMA), methacrylated gelatin (GelMA), gelatin, or a mixture thereof. Alternatively, the coating polymer may include a polymer formed by using surgical glue. Alternatively, the coating polymer may comprise gelatin. Alternatively, the coating polymer may include a polymer formed by using cyanoacrylate glue, fibrin sealant, collagen-based compound, glutaraldehyde glue, hydrogel, or a mixture thereof. Alternatively, the coating polymer may include a polymer formed by using n-butyl cyanoacrylate, 2-octyl cyanoacrylate, or a mixture thereof.

In the present disclosure, the mechanical strength of the ceramic scaffold may be in the range of 15N to 33N when measured as the maximum load of the stress-strain curve. And/or the bending strength of the ceramic scaffold may be in the range of 10MPa to 50 MPa; where the bending strength is measured by a standard three-point bending test. Alternatively, the mechanical strength of the ceramic support may be at least 5 times higher than the mechanical strength of the frame. Alternatively, the mechanical strength of the ceramic support may be at least 10 times greater than the mechanical strength of the frame. Alternatively, the mechanical strength of the ceramic support may be 10 to 20 times higher than that of the frame.

The present disclosure also relates to methods of making ceramic stents. The method can comprise the following steps: preparing a slurry comprising Hydroxyapatite (HA), tricalcium phosphate (TCP), or a mixture thereof; and a UV polymerizable monomer formulation; preparing a green body using a three-dimensional (3D) printing method and a slurry; debinding the green body and sintering to remove polymer formed by polymerization of the UV polymerizable monomer formulation to produce a sintered porous body, wherein the sintered porous body forms a framework; coating the sintered porous body with a polymer coating solution; polymerizing the polymer coating solution to form a coating comprising a polymer ("coating polymer"); thereby obtaining a ceramic support.

In the present disclosure, the ceramic scaffold may be sintered at a temperature in the range of 1000 to 1500 degrees celsius. Or wherein the green body may be sintered at a temperature in a range of 1050 degrees celsius to 1250 degrees celsius.

Bone defects are one of the most difficult injuries to heal that afflict the medical community for many years and seriously affect the quality of life of patients. Millions of surgeries are performed each year by using traditional gold standard implants, where significant limitations in size and osteogenesis remain prevalent. Three-dimensional (3D) scaffold-based therapies offer a promising solution, i.e. new bone can be regenerated with the help of osteoblasts. The mechanical strength and degradation of the scaffold play a key role in bone tissue regeneration. However, there are still significant challenges in the development of bone regeneration scaffolds due to the tradeoff between mechanical properties and degradation. A new 3D printing integration mixing method is researched, so that the mechanical performance is effectively improved by hundreds of times, and the degradation speed required by a 3D printing support is realized. Experiments aim at studying the influence of material selection and method design on the mechanical properties and degradation properties of the bioceramic scaffold. By using this proposed method a specific 3D scaffold designed for long bone and craniofacial circular defects was designed and manufactured. Based on our newly developed manufacturing method, a 3D trimmable plate was further invented to create a versatile and low cost solution for bone tissue regeneration. This approach opens up interesting prospects in the treatment of bone defects, with the potential for integration into expandable manufacturing.

To address these challenges, a 3D printing method with enhanced coatings was developed to fabricate bioceramic stents (fig. 1). 1 first, a green body of HA/TCP scaffold was prepared using paste printing based on mask image projection (MIP-SP). In some examples, the stent is printed in successive layers, each layer being set to a thickness of 10 μm to 200 μm, for example 10 μm, 20 μm, 30 μm, 40 μm, 50 μm, 60 μm, 70 μm, 80 μm, 90 μm, 100 μm, 120 μm, 140 μm, 160 μm, 180 μm or 200 μm. The HA/TCP scaffold is then degreased and sintered to remove the internal polymers and fuse the HA/TCP particles together, creating a micro-scale porous structure inside the scaffold. Then a biocompatible material with a faster degradation rate than HA/TCP is coated inside the HA/TCP stent, and the coating depth can be controlled by adjusting the values of different coating parameters. As the biodegradable coating material penetrates the HA/TCP stent and fills the empty spaces between the HA/TCP particles, the mechanical properties of the HA/TCP stent are significantly improved. By combining the coating method with MIP-SP, higher mechanical strength can be obtained at lower sintering temperature, so that HA/TCP stent can be degraded more quickly. It opens up a new way to achieve both mechanical properties and rapid degradation.

To solve the problem of reinforcement of HA/TCP stents, we first determined the effect of the coating method on the mechanical properties and degradation of HA/TCP stents. Therefore, we identified an enhanced mechanism for the coating process in HA/TCP stent fabrication. Later, material selection and process parameters were studied and optimized to change the properties of the HA/TCP scaffold. Two types of HA/TCP scaffolds were designed and manufactured for severe bone defects, including long bones and craniofacial bones, to demonstrate the feasibility of our 3D printing and augmentation method.

The ceramic scaffolds disclosed herein may be used for any type of bone replacement, including but not limited to the following types of bones: vertebrae, including cervical, thoracic, lumbar, sacral, caudal/vertebral (cordial); chest (thorax), hyoid, sternum, ribs; the skull comprises middle ear bone, skull comprises skull, occipital bone, parietal bone, frontal bone, temporal bone, sphenoid bone, ethmoid bone, facial bone, nasal bone, maxilla (maxilla), lacrimal bone, zygomatic bone (cheek bone), palatine bone, inferior turbinate, plough bone, mandible, middle ear, malleus, incus, and stapes; the arm bone includes: humerus, shoulder girdle (shoulder), scapula, clavicle, ulna, radius; hand bones including carpal bones, scaphoid bones, lunate bones, triquetrum bones, pisiform bones, trapezium bones, capitate bones, hamate bones, metacarpal bones, phalanges, proximal phalanges, middle phalanges, and distal phalanges; pelvis (pelvic girdle), ilium, ischium, and pubic bone; bones of the legs, including the femur, patella or patella, tibia, fibula; the metatarsals include tarsal (tarsus)/tarsal (tarsals), calcaneus or calcaneus, talus, navicular, medial cuneiform, lateral cuneiform, cuboid, metatarsal, phalanx, proximal phalanx, medial phalanx, and distal phalanx.

In addition, HA/TCP scaffolds have been designed for the treatment and study of bone defects, which opens up interesting prospects for clinical trials. However, efficient large-scale methods of designing and manufacturing HA/TCP scaffolds for bone defects are still lacking. With the benefit of the coating method, a 3D printed tailorable HA/TCP stent was designed and manufactured using the method we developed. It provides a new approach to fill the gap between customization and mass production. The method provided by the inventor solves the limitation of the current HA/TCP bracket, and HAs the feasibility of manual operation aiming at different purposes. Biological tests, including biocompatibility and cell attachment tests, were performed to demonstrate the greater operability and transformation potential of our proposed method in the treatment of large-scale bone defects.

Compared with other types of biological materials, the bioceramic HA/TCP HAs the advantages and advantages of cell attachment and proliferation, bone conduction, osseointegration and osteoinduction. However, the poor mechanical properties of HA/TCP scaffolds limit their widespread use in bone tissue regeneration. In this work, a new 3D printing method with a reinforced coating was proposed, so that the mechanical properties of the 3D printed HA/TCP stent were significantly improved without affecting the degradation of the HA/TCP stent. The impact of the reinforcement method on the mechanical properties and degradation of the stent is adjusted by the choice of coating material and control of the coating duration and process. Two types of post-coating methods, including surface spray and vacuum fusion, were applied to achieve different levels of mechanical reinforcement. The mechanism of coating enhancement is determined and verified through theoretical and experimental analysis. HA/TCP stents with special coatings show significant mechanical properties and degradation compared to traditional methods. We developed methods to design and manufacture 3DHA/TCP scaffolds for long bone and craniofacial severe bone defects to assess degradation and mechanical properties. In addition, a generic HA/TCP tailorable scaffold was constructed, providing the possibility of large scale manufacturing of HA/TCP scaffolds for general bone regeneration purposes. This opens up an attractive prospect for greater maneuverability and transformation potential for designing HA/TCP scaffolds based on enhanced coating methods to form HA/TCP scaffolds, thereby eliminating the bottleneck in current tissue engineering applications.

EXAMPLE 13 MIP-SP of HA/TCP scaffolds

The fabrication of bioceramic scaffolds with controlled micro-scale structural distribution is challenging using traditional fabrication methods such as freeze casting, foam replication, particulate leaching, and injection molding. Additive manufacturing techniques present advantages in building custom 3D scaffolds from scratch using different kinds of biomaterials. In this work, green bodies of3D HA/TCP scaffolds were first prepared using the self-developed MIP-SP method. In 3D printing of the green body, the digital model of the 3D scaffold was sliced to obtain a set of two-dimensional (2D) mask images with a 75 μm layer thickness, which was determined by the curing depth of the HA/TCP slurry (fig. 1, c). The generated mask image is used to instruct the digital micromirror to project light having a 2D pattern and the gray scale of each pixel in the mask image is adjusted to further control the light intensity. A uniform beam of light penetrates the transparent glass disc from the bottom and is focused on the top surface. Upon receiving sufficient energy from the exposure of the 2D patterned beam, a layer of HA/TCP slurry selectively transfers from the liquid phase to the solid phase due to the crosslinking reaction of the photocurable polymer, resulting in the encapsulation of HA/TCP particles inside the cured layer. After one layer is built, the platform is moved up and a new layer of HA/TCP slurry is fed in the light projection area using a blade assist material feed system as the transparent disk rotates. The platform is then moved downward to form a uniform layer of slurry for fabrication of the next layer. Following the above procedure, the 2D patterned HA/TCP slurry was stacked layer by layer and finally the viscous HA/TCP slurry was formed into the desired 3D shape.

After the 3D printing process, green HA/TCP scaffolds were prepared. Post-processing, including degreasing and sintering, is then performed to remove the internal polymer and fuse the HA/TCP particles [33 ]. After degreasing, the HA/TCP particles are loosely arranged and the internal porous structure is partially reduced during sintering at much higher temperature settings. During sintering, point contacts between HA/TCP particles become grain boundaries, and as the sintering temperature increases, the grains become larger, resulting in decreased porosity and shrinkage of the HA/TCP scaffold. The porosity of the HA/TCP scaffold was adjusted by varying the sintering temperature and the mass concentration of HA/TCP particles. For example, the shrinkage ratio and porosity of a 30% HA/TCP scaffold sintered at 1050 ℃ are 17.75% and 20%, respectively, and the shrinkage ratio of the HA/TCP scaffold increases with increasing sintering temperature (see fig. 3, g). Based on the shrinkage ratio of the HA/TCP scaffold after the sintering process, a compensation operation is performed to obtain the desired shape of the HA/TCP scaffold. In addition, the mechanical properties of the HA/TCP scaffold are improved due to the grain growth. However, the compressive strength of HA/TCP scaffolds is only a few MPa to a dozen MPa when sintered at temperatures ranging from 1050 to 1250 degrees. The 3D printed HA/TCP stent is too fragile to withstand any further manipulation in subsequent surgery. Thus, the poor mechanical properties of the bioceramic scaffold limit its application in biological tissue regeneration. In this work, the second stage coating process was integrated into the manufacture of HA/TCP stents to significantly improve their mechanical properties. Meanwhile, the biodegradability of the HA/TCP stent remains unchanged. The difference in coating material and coating material penetration depth results in widely tunable mechanical and biodegradable properties.

Example 2 mechanical Property characterization

The mechanical properties of the 3D printed HA/TCP stent with/without coating material were evaluated by the maximum load that the stent can withstand, which indicates the compressive strength of the 3D printed HA/TCP component. Figure 2 shows that the maximum load of the HA/TCP microcell printing element without coating material is only 0.04N due to poor bonding between the HA/TCP particles. However, the coating method is used to gradually fill the coating material inside the pores of the 3D printed HA/TCP part to enhance the binding of the HA/TCP particles. We first evaluated the mechanical properties of HA/TCP after coating treatment by using different coating materials. After surface coating, the maximum load of the HA/TCP microcell printing components of surgical glue and PEGDMA increased to 2.8N and 2.2N, respectively. More specifically, cyanoacrylate-based polymers in surgical glues will polymerize by an anionic mechanism, eventually forming long-chain polymers. Following the coating process, the mechanical properties of the 3D printed HA/TCP parts are effectively enhanced with strong bonding between the polymer and the HA/TCP particles. Other types of coating materials, such as acrylate based photo-curable polymers and gelatin, will react with different crosslinking mechanisms, which help to enhance the binding of the HA/TCP particles, which determines the mechanical properties of the HA/TCP printed part (fig. 2, c-d). The evolution of the bond during the coating process was verified by Scanning Electron Microscope (SEM) images. During the coating process, the coating solution gradually penetrates inside the 3D printed HA/TCP part with micron-sized pores. After the coating process, the polymer in the coating solution is crosslinked such that the HA/TCP particles are trapped in the network of coating material. As shown in fig. 2, c, a thin polymer coating is produced and mixing with the underlying HA/TCP particles is evident, indicating intimate bonding between the coating material and the HA/TCP particles. It was noted that even when the penetration depth of the coating material was relatively small (thickness only hundreds of micrometers), the maximum load of the 3D printed HA/TCP part with acrylate based coating material had increased more than 100 times the maximum load of the original HA/TCP printed part without coating. By exploring coating materials with high mechanical properties and increasing the penetration depth of the coating material, stronger binding between the HA/TCP particles and the coating material can be obtained to further enhance the mechanical properties of the HA/TCP stent. Since the coating material we choose is a rapidly degradable material without toxic components, this makes future biological applications possible.

By varying the permeability of the coating material, a wide range of tunable mechanical and biodegradable properties can be achieved. The effect of coating depth on the mechanical properties of HA/TCP printed parts was evaluated. Both surface spray coating and vacuum fusion are applied in the coating process to achieve various penetration depths of the coating material inside the HA/TCP printed part (see fig. 3, a). For both coating methods, the maximum load and fracture toughness increased with increasing penetration depth.

(1) In surface spraying, coating thicknesses ranging from several micrometers to several hundred micrometers are obtained, for example, 5 μm, 10 μm, 20 μm, 30 μm, 40 μm, 50 μm, 60 μm, 70 μm, 80 μm, 90 μm, 100 μm, 150 μm, 200 μm, 250 μm, 300 μm, 400 μm, 500 μm, 600 μm, 700 μm, 800 μm, 900 μm or 1mm, and the maximum load and fracture toughness (i.e., flexural strength) of the HA/TCP printed part having the maximum coating layer thickness by using surface spraying are 8N and 0.16MPa · m, respectively ^1/2 (see FIG. 3, c-f). The difference in mechanical properties of HA/TCP printed parts with one type of coating material mainly results from the penetration depth of the coating material. In addition to surface spraying, brush spraying may also be used to apply the coating material to certain areas of the stent. Similar coating thicknesses ranging from a few microns to a few hundred microns were obtained and the mechanical properties of the coated HA/TCP printed parts were similar to those of the parts with surface spray coating as shown in fig. 3, e-f.

(2) In addition, vacuum fusion can significantly increase the penetration depth of the coating material in the 3D printed stent. In some examples, the penetration depth is 50 μm, 100 μm, 150 μm, 200 μm, 250 μm, 300 μm, 350 μm, 400 μm, 450 μm, 500 μm, 550 μm, 600 μm, 700 μm, 800 μm, 900 μm, 1mm, 1.25mm, 1.5mm, 2mm, 2.25mm, 2.5mm, 2.75mm, 3mm, 3.25mm, 3.75mm, 4mm, 4.25mm, 4.75mm, or 5mm, or any range between two of the foregoing values. More specifically, after the sintering process, micro-and nano-scale pores are generated inside the HA/TCP stent (fig. 3, b), and the coating material filled in the pores of the HA/TCP printing part is sucked out by the vacuum generator together with the internal air. As shown in fig. 3, d, the penetration depth of the coating material is varied with the coating timeAnd increases with increasing time. For example, it takes 20s and 35s to coat surgical glue with a penetration depth of 500 μm and 1.4 mm. After full application of the surgical glue, the maximum load and fracture toughness of the HA/TCP printed part increased to 19N and 0.40MPa.m, respectively ^1/2 (FIG. 3, c-f). The maximum load of the HA/TCP printing component is determined more by the coating method. The maximum load and fracture toughness were improved more than 450 and 26 times, respectively, compared to pure HA/TCP printed parts without a coating process.

In addition to the change in mechanical properties, the shape of the HA/TCP printed part also changed after the sintering and coating process. Specifically, the HA/TCP printed components shrink after the sintering process because the internal gaps between the HA/TCP printed components decrease as the grains grow. Shrinkage ratio R of HA/TCP printing component _s Increases with increasing sintering temperature. The shrinkage of the HA/TCP printing element is not uniform, wherein the shrinkage ratio of the HA/TCP printing element in the axial direction is greater than the shrinkage ratio thereof in the radial direction. The shrinkage results of the HA/TCP printed parts sintered at different temperatures are shown in figure 3, g. However, HA/TCP printed parts expand after the coating material fills the internal pores. Similarly, the expansion of the HA/TCP printing element exhibits anisotropy in the axial and radial directions. The expansion ratio of the HA/TCP printing element in the axial direction is greater than its expansion ratio in the radial direction. The expansion ratio R increases with the sintering temperature _e Lower as HA/TCP printed parts become more dense with smaller porous structures. According to the experimental results, the size compensation of the printed model needs to be applied to the input CAD model to accurately control the shape of the HA/TCP stent. Adjusting the size of the HA/TCP stent according to the shape change rate:

(1-r _s )(1+r _e )，

and the compensation coefficient phi can be calculated by:

furthermore, the sintering temperature has an influence on the mechanical and biodegradable properties. The compressive strength of HA/TCP scaffolds with microcellular structures sintered at different temperatures were compared (see FIG. 4). The cracked brittle fracture occurred first in the HA/TCP stent region without coating material (fig. 4, d) and gradually led to catastrophic failure (fig. 4, b). However, the HA/TCP printed part with the coating material did not fail much and there was only a small fracture at the same load compared to the HA/TCP printed part without the coating material (fig. 4, c). Structural simulations using COMSOL Multiphysics showed that stresses are concentrated at the junction areas between HA/TCP particles (fig. 4, b-c). The uneven stress distribution can easily break the grain boundaries between the HA/TCP particles and lead to cracks branching along the distribution of the HA/TCP particles without coating (fig. 4 d).

The stress-strain characteristics of3D HA/TCP printed parts with coated materials sintered at different temperatures were investigated. As the sintering temperature increased, the grain boundaries became larger, and the simulation results showed that stress was concentrated at the grain boundaries of the HA/TCP particles (fig. 4, e). The stress-strain curves of the HA/TCP microcell structures sintered at 1250 deg.c with and without the surgical glue coating are shown in fig. 4, f. FIG. 4, g shows the maximum load of HA/TCP scaffolds with microcellular structure sintered at different temperatures (including 1050 deg.C, 1150 deg.C and 1250 deg.C). The maximum load of the HA/TCP printing component decreases with decreasing sintering temperature. For example, HA/TCP sintered at 1050 ℃ and 1250 ℃ have maximum loads of 15N and 33N, respectively. As the sintering temperature increases, the grains of the HA/TCP particles become larger, thereby enabling them to withstand greater stresses.

In the case of a fully coated material HA/TCP printed part, the nucleophile initiates a reaction with the hydroxide in HA by attacking the carbon-carbon covalent bond, which breaks and forms a new bond. After polymerization of the cyanoacrylate-based coating material, a strong bonding network is formed between the HA/TCP particles and the cyanoacrylate. The enhanced binding of the polymer to the HA/TCP particles improves the mechanical properties of the material. The loading was further increased to disrupt the bond between the polymer matrix and the HA/TCP particles. Thus, the maximum load of the coated HA/TCP stent sintered at 1250 ℃ is greater than the maximum load of the coated HA/TCP stent sintered at 1050 ℃. However, the coating of the HA/TCP stent sintered at 1250 ℃ did not improve the compressive strength significantly as the stent sintered at 1050 ℃.

The effect of the coating enhancement method on the flexural strength and fracture toughness was determined by performing a standard three-point bending test. 3D printed HA/TCP components (77.14mm x 1.07mm x1.42mm) with and without a surgical glue coating were tested to determine the reinforcing mechanism of the coating material (FIG. 5, a). Fig. 5, b-c show the load-displacement curve and fracture toughness of the HA/TCP printed part before and after the coating process of sintering at 1050 ℃, respectively. For example, the fracture toughness of HA/TCP printed parts sintered at 1150 ℃ is improved by 18 times from only 0.015MPa ^1/2 To 0.27MPa.m ^1/2 . COMSOL Multiphysics simulation results show that damage in pure HA/TCP is only located at grain boundaries, which exhibits unstable cracking characteristics; in contrast, the addition of the polymer stopped the main cracks and deflected the micro cracks (see fig. 5, e). The results show that pure HA/TCP ceramic printing components quickly produced catastrophic failure, with flexural cracks occurring within short bending distances (fig. 5, f). After coating the polymer-based material, the cyanoacrylate-based polymer still showed the ability to hold all crack parts together, even though there were slight cracks in the grain boundaries between the HA/TCP particles. As shown in fig. 5, g, cracks deflected and occurred at the interface between HA/TCP and cyanoacrylate polymer, and the cracks were bridged by macrofibers and microfibers, resulting in a significant increase in fracture toughness.

Sintering temperature also affects fracture toughness and flexural strength. For example, after the coating process, the bending strength of the HA/TCP printed component sintered at 1050 ℃ is twice that of the HA/TCP printed component sintered at 1250 ℃. Furthermore, the flexural strength of the HA/TCP printed part sintered at 1050 ℃ was improved 260 times over the HA/TCP printed part without the coating material. The coating process significantly improves the bending strength and fracture toughness of the 3D printed HA/TCP part because the coating material plays a dominant role in the bending performance of the 3D printed HA/TCP part.

The bridged polymeric macrofibers and polymeric microfibers also contribute to improved tensile strength of the HA/TCP printed part. As shown in fig. 5, i, the HA/TCP printed component sintered at 1050 ℃ HAs a greater displacement than the HA/TCP printed component sintered at 1250 ℃, because there is more coating material inside the printed component. Due to the addition of the polymer, the young's modulus of the HA/TCP printing component increased by a factor of one hundred. Pure HA/TCP exhibits brittle fracture under small tensile forces and catastrophic failure under light loads. Unlike the pure HA/TCP polymer, the polymer microfibers were evident on both sides of failure, forming saw-tooth like breaks. Similarly, the difference in young's modulus of the HA/TCP printed part after the coating process sintered at different temperatures is attributed to the proportion of coating material inside the HA/TCP printed part. Due to the dense HA/TCP particles with less polymer coating, brittle fracture occurred in HA/TCP printed parts sintered at 1250 ℃ due to crack formation and rapid propagation. The coated HA/TCP printed part sintered at 1050 ℃ showed ductility because the coating material surrounding the HA/TCP particles was connected (see fig. 5, k).

In addition to the surgical glue we extensively studied, gelatin as a coating material showed significant mechanical property improvements and desirable surface properties with micropores for cell attachment (see fig. 2, c-d). The coating material may also be diluted with a percentage of water prior to the coating process. This may result in increased biodegradability of the 3D printed HA/TCP components.

Example 3 printing a modifiable plate for a universal craniofacial bone defect

We have produced HA/TCP reformable plates using our proposed method for the general purpose of bone defects. The CAD model of the HA/TCP trimmable board is shown in FIG. 6, a. The overall height of the green HA/TCP was 2mm and the overall XY dimension of the holder was 40mm by 40mm, the holder being decorated with a 2.5mm array of through holes. To obtain better surface quality, the slice thickness at the scaffold location was set to 75 μm per layer and the grey level of the 2D patterned beam was varied based on the curing performance database described previously. However, any size stent may be manufactured and any slice thickness of the stent layer may be used to print the stent. FIG. 6, b shows the printed results of the HA/TCP board before degreasing and sintering. The sintered and coated HA/TCP boards are shown in FIGS. 6, c-d, respectively. Due to the enhancement of the coating process, the HA/TCP plate can be cut into the desired shape to accommodate the needs of different bone defects. Scanning Electron Microscopy (SEM) of the HA/TCP plate with the coating material produced is shown in fig. 6, g. The array of through holes is designed to mount the screws at the fracture site and the fixation position can be adjusted according to the actual surgical situation. It can be seen that the surface quality of the HA/TCP trimmable boards produced with the enhanced coating is satisfactory. Furthermore, various scaffold plates designed to different thicknesses and sizes can be manufactured using our proposed method, enabling large-scale preparation of HA/TCP scaffolds for universal bone reconstruction. In some examples, the scaffold thickness may be in the range of 100 μm to 100mm, for example 100 μm, 200 μm, 300 μm, 400 μm, 500 μm, 600 μm, 700 μm, 800 μm, 900 μm, 1mm, 1.2mm, 1.4mm, 1.6mm, 1.8mm, 2mm, 3mm, 5mm, 10mm, 15mm, 20mm, 25mm, 30mm, 35mm, 40mm, 45mm, 50mm, 60mm, 70mm, 80mm, 90mm, or 100 mm.

Example 4 3D printed scaffold for craniofacial gross defects

Another craniofacial test case we tested was to construct a 3DHA/TCP scaffold for severe bone defects (as shown in FIG. 7, a). FIG. 7, b shows a CAD model of HA/TCP scaffold composed of microcell structures. The height and diameter of the HA/TCP stent were 8mm and 40mm, respectively. First, green HA/TCP scaffolds were fabricated using MIP-SP. Then fig. 7, d shows the pure HA/TCP scaffold after degreasing and sintering. High strength HA/TCP stents were then fabricated using the reinforced coating method (see fig. 7, e). The compression simulation of a pure HA/TCP stent using COMSOL is shown in FIG. 7, f, where the joint region is subjected to large stresses. The relevant compression tests of HA/TCP stents with and without the reinforced coating process were performed (see fig. 7, g (without reinforced coating), fig. 7, h (with reinforced coating)). The pure HA/TCP scaffold was crushed after compression with only a few newton forces. However, HA/TCP stents with coated materials are able to retain shape after compression. Until the compressive load was gradually increased to 55N, only small fractures were observed in the HA/TCP stent with the coated material. Fig. 7, i shows the details of the force and displacement of the HA/TCP stent with and without coating.

Example 5 regeneration of femoral severe defects

As another example, a thin shell-like HA/TCP scaffold for a mouse model of severe defect of femur can also be constructed (see fig. 8, a). The CAD model of the HA/TCP scaffold is shown in FIG. 8, b, and the thickness and height of the scaffold are 100 μm and 2mm, respectively. Micro-scale holes of 150 μm are added on the side wall of the scaffold for transporting nutrients. First, a MIP-SP was used to prepare HA/TCP scaffolds (FIG. 8, c). Degreasing and sintering were then performed to remove the internal polymer, resulting in a pure HA/TCP scaffold, as shown in fig. 8, d. The COMSOL software system was used to simulate the stress and strain of HA/TCP under compression (FIG. 8, e). To improve the mechanical properties, the HA/TCP stent is filled with a coating material inside, and fig. 8, f shows a view of the HA/TCP stent manufactured with the coating. Furthermore, we investigated the compressive properties of the coated and uncoated HA/TCP stents. The maximum compressive load of the HA/TCP scaffold was increased nearly ten times compared to the original pure HA/TCP sintered at different temperatures (see figure 8, g). The mechanical properties of the stent meet the experimental requirements and there is no HA/TCP stent failure during the surgical procedure.

Example 6 HA/TCP slurry

Micro-scale Hydroxyapatite (HA) and tricalcium phosphate (TCP) powders were purchased from Sigma-Aldrich and used to make HA/TCP scaffolds. The average diameters of HA and TCP powders were 10 μm and 4 μm, respectively. The polymeric binder to be removed after the degreasing process is necessary to form the HA/TCP powder into a 3D shape during the 3D printing process. To prepare an HA/TCP based slurry, 15 wt% HA and 15 wt% TCP were first mixed into a photocurable liquid resin WaxPast, purchased from MakerJuise Labs. The components of the liquid resin are acrylates, photoinitiators, crosslinkers and stabilizers, and the photocurable polymer can crosslink under exposure to visible light. The HA/TCP suspension was then ball milled for 40 minutes at 200rpm to ensure uniform distribution of the HA/TCP powder within the photocurable resin. Finally, residual gases in the HA/TCP slurry were removed by evacuation with a special apparatus for later use.

Example 7 MIP-SP

A MIP-SP prototype consisting of an optical module, a material supply module and a motion module was constructed for the manufacture of green HA/TCP scaffolds. In thatIn optical modules, a Digital Micromirror (DMD) device (Texas instruments) has millions of micromirrors that can be individually controlled to set their states on or off [ 38)]. In the on state, light is reflected so that the pixels in the projection area appear bright on the top surface of the transparent disk. The brightness of each pixel in the 2D patterned beam can be adjusted by controlling the angle of the micromirror. The fabrication area of MIP-SP was 106X 60mm in our prototype ² And the resolution of the curing beam is 55 microns per pixel. Since the viscosity of the 30 wt% HA/TCP slurry was 5000mPa-s, this material was unable to refill back into the printed area when driven by air pressure and gravity alone. The refilling problem of HA/TCP slurry is solved by adopting a scraper auxiliary material feeding system. A thin layer of HA/TCP slurry is formed by a blade moving along a transparent glass plate and the thickness of the material can be adjusted by varying the speed of movement and the gap distance between the blade and the transparent frame. To speed up the slurry coating process, the transparent disk was mounted on a rotating table and a 100 μm thick layer was continuously coated on the transparent disk. All HA/TCP parts studied in this work were printed using MIP-SP and the exposure time for each layer cure was set to 10 seconds.

Example 8 degreasing and sintering

The degreasing process is performed under a vacuum environment using a tube furnace to smoothly remove the internal polymer of the HA/TCP printed part and to avoid cracks when the decomposition rate of the polymer exceeds its pyrolysis rate. The generated gas was continuously sucked out, and the internal pressure of the heating zone was maintained at-0.1 MPa. Specifically, in the degreasing process, the heating rate was set to 1 degree/min, and the temperature was maintained at 500 ° F, 900 ° F, 1200 ° F for 2 hours, respectively. The temperature was then cooled to 1000 ° F for one hour to completely burn off the internal polymer. And finally, naturally cooling the temperature to room temperature. After the degreasing process, the HA/TCP powder in the printed part is sparsely arranged, resulting in poor mechanical properties. To solve this problem, an additional sintering process must be performed after the degreasing process.

During sintering, the temperature setting is much higher so that the HA/TCP particles can grow and fuse together. The shrinkage rate and the mechanical property of the HA/TCP support can be adjusted by changing the sintering temperature. The HA/TCP scaffolds were sintered in a tube furnace under normal air conditions. To obtain a wide range of mechanical properties, the sintering temperature is set to three different levels. For all temperature settings, the ramp rate was set to 5 degrees/min and the temperature was maintained at 300 ℃, 600 ℃ and 900 ℃. After that, the temperature was raised to 1050 ℃, 1150 ℃ and 1250 ℃, respectively, and further kept at the peak value for 3 hours [33 ]. And finishing the sintering process after natural cooling.

Example 9 coating materials and methods

Surgical glues, internal tissue glues and burning matrigel matrices (recon base films) available from Gluustich, COHERA medical and Careford, respectively, may be used. The components of the surgical glue are n-butyl cyanoacrylate and 2-octyl cyanoacrylate.

A100% (w/v) poly (ethylene glycol) dimethacrylate (PEGDMA, Mw750, Sigma-Aldrich) solution was prepared according to the following procedure: 1% (w/v) visible light photoinitiator (Irgacure 819, BASF) was first dissolved completely in Phosphate Buffered Saline (PBS) to initiate chain polymerization by free radicals, and then 100% (w/v) poly (ethylene glycol) dimethacrylate (PEGDMA, Mw750, Sigma-Aldrich) was mixed in solution with magnetic stirring.

15% (w/v) methacrylated gelatin (GelMA): gelatin from pig skin (Sigma-Aldrich) was first mixed at 10% (w/v) with Dulbecco's phosphate buffered saline (DPBS; Gibco) by gently swirling the mixture in a water bath at 50 ℃ for 15 minutes and stirring until complete dissolution. A high degree of methacrylation was achieved by adding 20% (w/v) methacrylic anhydride (MA, Sigma-Aldrich) to the synthesis reaction at a rate of 0.5mL/min with stirring at 50 ℃ and allowing to react for 2 hours. After that, 5-fold dilution with DPBS was added to stop the reaction, and the mixture was dialyzed with distilled water at 40 ℃ for 1 week using a 12-14 kDa cut-off dialysis tube to remove salts and residual MA. The 0.2% gelatin solution was cooled at room temperature, heated at about 37-40 ℃ and filtered through a 0.45 μm cellulose acetate membrane (CA). The filtered solution was lyophilized for 1 week to give a white porous foam and stored at-80 ℃. Then 1% (w/v) Irgacure 819 was completely dissolved in Phosphate Buffered Saline (PBS) to induce chain polymerization by free radicals, and 15% (w/v) GelMA was gradually added to the solution under magnetic stirring until GelMA was completely dissolved in the solution.

Gelatin solution: 0.25qt of water was boiled at 100 ℃ and 7g of gelatin powder (Knox) was sprinkled in the water. The mixture was allowed to stand for 1 minute and stirred for 5 minutes until the gelatin powder was completely dissolved. All of the above solutions were degassed in vacuo prior to the coating process.

Each of the above coating polymer solutions may be used alone or in combination as a mixture of two or more polymer solutions.

In surface spraying, a conventional sprayer is used to force the coating material through a nozzle that disperses the flow of coating material through the use of a one-way valve. Micron-sized mist is generated and further deposited on the surface of the printed HA/TCP part. The thickness of the deposited layer is mainly determined by the spraying time. The coating material only covers the surface by the spray process and the coating is difficult to penetrate the HA/TCP printed part, resulting in relatively limited improvement of mechanical properties. To further improve mechanical properties, this work investigated vacuum fusion. In vacuum fusion, the printed HA/TCP printing components are incorporated into a reservoir filled with coating material, and the entire reservoir is placed in a vacuum environment. The vacuum conditions and coating time are adjusted to control the penetration depth of the coating material. For example, when the air pressure was set to 25in · Hg, the coating speed of the surgical glue was 20 μm/s for the HA/TCP printed part sintered at 1150 ℃.

EXAMPLE 10 mechanical testing

A series of tests including compression, three point bending deflection, fracture toughness and tensile tests were performed using a universal Testing machine (Instron 5492Dual Column Testing Systems, Instron, MA, USA). The mechanical properties of each experimental group were evaluated using three printing units. For the compression test, a static compression model with a compression speed of 5mm/min and a maximum compression distance of 2mm was selected for the test. The 3D printed test printed component was a separate unit from the integrated printed component having a height of 5mm, a width of 2mm and a mesh thickness of 0.7 mm. After sintering, the green body was then placed vertically in the middle of the test platform for compression testing. The strength and strain were calculated using the following equations:

where F is the load, r is the diameter, and L0 is the height of the 3D printing component.

For the three point bend flexural test, rectangular parallelepiped printing elements were printed with dimensions of 7.14mm long, 1.07mm wide and 1.42mm high. After sintering, the parts were printed for a three point bend flexural test. The sample was placed on a printing part holder spanning 5 mm. The test load was applied at the midpoint of the sample at a rate of 5 mm/min. The bending strength was calculated by:

where F is the force at the break point, L is the span, b is the print part width, and d is the part thickness [40 ].

To determine the fracture toughness of3D printed parts, we used a Single Edge Notch Bending (SENB) test. The printing part was identical to the printing part of the three-point bending deflection test except that it had a notch on one side. The sample was placed on a 3D printing part holder spanning 5 mm. The test load is applied at the midpoint of the specimen, i.e., the top of the notch. The test was carried out at a loading rate of 5 mm/min. Fracture toughness is calculated by using the following equation:

where P is the maximum load during the SENB test, S is the support span, b is the print part width, d is the sample thickness, and a is the notch depth.

The tensile test was a 3D printed sample 9.5mm long by 2.5mm wide by 1mm thick. The tensile test was carried out at a loading rate of 5 mm/min. And the elastic modulus is obtained by calculating the slope of the linear region on the stress-strain curve. The stress and strain are respectively calculated by using the following equations.

Where F is the load, b is the width, D is the length, and L0 is the height of the 3D printing component.

Statistical analysis

For each statistical analysis, experiments were performed by using R statistics software. All data are expressed as mean ± Standard Deviation (SD). The significance parameter in each trial was determined using one-way analysis of variance (ANOVA), and statistical significance was considered as p < 0.05.

All articles, patents, patent applications, and other publications cited in this disclosure are incorporated herein by reference.

In this disclosure, the indefinite articles "a" and "an" are synonymous with the phrases "one or more" and "at least one", and mean "at least one".

Relational terms such as "first" and "second," and the like may be used solely to distinguish one entity or action from another entity or action without necessarily requiring or implying any actual such relationship or order between such entities or actions. The terms "comprises," "comprising," and any other variations thereof, when used in conjunction with a list of elements in the specification or claims, are intended to indicate that the list is not exclusive and may include other elements. Similarly, without further limitation, elements prefaced by the word "a" or "an" do not exclude the presence of other elements of the same type.

The abstract is provided to assist the reader in quickly determining the nature of the technical disclosure. It is submitted with the understanding that it will not be used to interpret or limit the scope or meaning of the claims. In addition, various features in the foregoing detailed description are grouped together in various examples to simplify the present disclosure. This method of disclosure is not to be interpreted as reflecting an intention that the claimed examples require more features than are expressly recited in each claim. Rather, as the following claims reflect, inventive subject matter lies in less than all features of a single disclosed example. Thus the following claims are hereby incorporated into the detailed description, with each claim standing on its own as a separately claimed subject matter.

Reference to related art

The following publications are related art in the context of this disclosure.

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Claims

1. A ceramic stent, comprising:

a coating comprising a coating polymer;

wherein the frame has at least one surface, wherein the coating is formed on the at least one surface of the frame, and wherein the coating at least partially covers the at least one surface of the frame.

2. The ceramic stent of claim 1, wherein the coating polymer comprises a polymer formed by using surgical glue, a colloidal protein mixture, poly (ethylene glycol) dimethacrylate (PEGDMA), methacrylated gelatin (GelMA), gelatin, or a mixture thereof.

3. The ceramic stent of claim 2, wherein the coating polymer comprises surgical glue, gelatin, or a mixture thereof.

4. The ceramic stent of claim 3, wherein the coating polymer comprises n-butyl cyanoacrylate monomer, 2-octyl cyanoacrylate monomer, or a mixture thereof.

5. The ceramic stent of claim 4, wherein the coating is coated on the surface of the stent at a thickness of 5 μm to 1 mm.

6. The ceramic stent of claim 5, wherein the mechanical strength of the ceramic stent is in the range of 15N to 33N when measured as the maximum load of a stress-strain curve.

7. The ceramic stent of claim 5, wherein the bending strength of the ceramic stent is in the range of 10MPa to 50 MPa; wherein the bending strength is measured by a standard three-point bending test.

8. The ceramic stent of claim 5, wherein the mechanical strength of the ceramic stent is 10 to 20 times higher than the mechanical strength of the frame.

9. The ceramic scaffold of any preceding claim, wherein the coating polymer comprises a polymer formed by using cyanoacrylate glue, fibrin sealant, collagen-based compounds, glutaraldehyde glue, hydrogels, or mixtures thereof.

10. The ceramic stent of any one of the preceding claims, wherein the coating polymer comprises an acrylate polymer.

11.The ceramic stent of any one of the preceding claims, wherein the coating polymer comprises a cyanoacrylate polymer.

12. The ceramic stent according to any one of the preceding claims, wherein the coating polymer comprises a surgical glue.

13. The ceramic stent according to any one of the preceding claims, wherein the coating polymer comprises gelatin.

14. The ceramic stent of any one of the preceding claims, wherein the coating polymer comprises a polymer of n-butyl cyanoacrylate monomer, 2-octyl cyanoacrylate monomer, or a mixture thereof.

15. The ceramic stent according to any one of the preceding claims, wherein the coating is coated on the surface of the stent at a thickness of 5 μm to 1 mm.

16. The ceramic scaffold according to any one of the preceding claims, wherein the scaffold comprises stacked layers of Hydroxyapatite (HA), tricalcium phosphate (TCP) or a mixture thereof, each layer having a thickness of 10 to 200 μm.

17. The ceramic stent according to any one of the preceding claims, wherein the mechanical strength of the ceramic stent is in the range of 15N to 33N when measured as the maximum load of a stress-strain curve.

18. The ceramic stent of any one of the preceding claims, wherein the ceramic stent has a flexural strength in the range of 10MPa to 50 MPa; wherein the flexural strength is measured by a standard three point bending test.

19. The ceramic scaffold of any of the preceding claims, wherein the mechanical strength of the ceramic scaffold is at least 5 times greater than the mechanical strength of the frame.

20. The ceramic scaffold of any of the preceding claims, wherein the mechanical strength of the ceramic scaffold is at least 10 times greater than the mechanical strength of the frame.

21. The ceramic scaffold of any of the preceding claims, wherein the mechanical strength of the ceramic scaffold is 10-20 times greater than the mechanical strength of the frame.

22. The ceramic stent according to any one of the preceding claims, wherein the coating polymer has a thickness in the range of 1 micron to 1000 microns.

23. The ceramic stent according to any one of the preceding claims, wherein the coating polymer has a thickness in the range of 10 to 500 microns.

24. The ceramic stent of any one of the preceding claims, wherein the coating polymer comprises a polymer formed by using surgical glue, a colloidal protein mixture, poly (ethylene glycol) dimethacrylate (PEGDMA), methacrylated gelatin (GelMA), gelatin, or a mixture thereof; wherein the mechanical strength of the ceramic support is in the range of 15N to 33N, and/or the flexural strength of the ceramic support is in the range of 10MPa to 50MPa, when measured as the maximum load of a stress-strain curve; and wherein the bending strength is measured by a standard three-point bending test.

25. The ceramic stent of any one of the preceding claims, wherein the coating polymer comprises a polymer formed by using surgical glue, a colloidal protein mixture, poly (ethylene glycol) dimethacrylate (PEGDMA), methacrylated gelatin (GelMA), gelatin, or a mixture thereof; wherein the mechanical strength of the ceramic scaffold is at least 5 times greater than the mechanical strength of the frame, or at least 10 times greater than the mechanical strength of the frame, or 10 to 20 times greater than the mechanical strength of the frame; and wherein the mechanical strength of the ceramic scaffold is the maximum load of the stress-strain curve.

26. The ceramic stent of any one of the preceding claims, wherein the coating polymer comprises a polymer formed by using n-butyl cyanoacrylate, 2-octyl cyanoacrylate, or a mixture thereof; wherein the mechanical strength of the ceramic support is in the range of 15N to 33N and/or the bending strength of the ceramic support is in the range of 10MPa to 50MPa when measured as the maximum load of a stress-strain curve; and wherein the bending strength is measured by a standard three point bending test.

27. The ceramic stent of any one of the preceding claims, wherein the coating polymer comprises a polymer formed by using n-butyl cyanoacrylate, 2-octyl cyanoacrylate, or a mixture thereof; wherein the mechanical strength of the ceramic scaffold is at least 5 times greater than the mechanical strength of the frame, or at least 10 times greater than the mechanical strength of the frame, or 10 to 20 times greater than the mechanical strength of the frame; and wherein the mechanical strength of the ceramic scaffold is the maximum load of the stress-strain curve.

28. The ceramic stent according to any one of the preceding claims, wherein the coating substantially covers the at least one surface of the frame.

29. The ceramic stent according to any one of the preceding or following claims, wherein said coating polymer comprises a polymer formed by using surgical glue, a colloidal protein mixture, poly (ethylene glycol) dimethacrylate (PEGDMA), methacrylated gelatin (GelMA), gelatin, or a mixture thereof.

30. The ceramic stent according to any one of the preceding or following claims, wherein the coating polymer comprises a polymer formed by using surgical glue.

31. The ceramic stent according to any one of the preceding or following claims, wherein the coating polymer comprises a polymer formed by using gelatin.

32. The ceramic scaffold of any preceding or subsequent claim, wherein the coating polymer comprises a polymer formed by using a cyanoacrylate glue, a fibrin sealant, a collagen-based compound, a glutaraldehyde glue, a hydrogel, or a mixture thereof.

33. The ceramic stent of any one of the preceding or following claims, wherein the coating polymer comprises a polymer formed by using n-butyl cyanoacrylate, 2-octyl cyanoacrylate, or a mixture thereof.

34. The ceramic stent according to any one of the preceding or following claims, wherein the mechanical strength of the ceramic stent is in the range of 15N to 33N when measured as the maximum load of a stress-strain curve.

35. The ceramic stent according to any one of the preceding or following claims, wherein the bending strength of the ceramic stent is in the range of 10MPa to 50 MPa; wherein the bending strength is measured by a standard three-point bending test.

36. The ceramic scaffold of any preceding or following claim, wherein the mechanical strength of the ceramic scaffold is at least 5 times greater than the mechanical strength of the frame.

37. The ceramic scaffold of any preceding or following claim, wherein the mechanical strength of the ceramic scaffold is at least 10 times greater than the mechanical strength of the frame.

38. The ceramic scaffold of any preceding or following claim, wherein the mechanical strength of the ceramic scaffold is 10 to 20 times greater than the mechanical strength of the frame.

39. A ceramic stent according to any one of the preceding or following claims, wherein the coating polymer has a thickness in the range of 1 micron to 1000 microns.

40. A ceramic stent according to any one of the preceding or following claims, wherein the coating polymer has a thickness in the range of 10 to 500 microns.

41. The ceramic stent of any one of the preceding or following claims, wherein the coating polymer comprises a polymer formed by using surgical glue, a colloidal protein mixture, poly (ethylene glycol) dimethacrylate (PEGDMA), methacrylated gelatin (GelMA), gelatin, or a mixture thereof; wherein the mechanical strength of the ceramic support is in the range of 15N to 33N and/or the bending strength of the ceramic support is in the range of 10MPa to 50MPa when measured as the maximum load of a stress-strain curve; and wherein the bending strength is measured by a standard three-point bending test.

42. The ceramic stent of any one of the preceding or following claims, wherein the coating polymer comprises a polymer formed by using surgical glue, a colloidal protein mixture, poly (ethylene glycol) dimethacrylate (PEGDMA), methacrylated gelatin (GelMA), gelatin, or a mixture thereof; wherein the mechanical strength of the ceramic scaffold is at least 5 times higher than the mechanical strength of the frame, or at least 10 times higher than the mechanical strength of the frame, or 10 to 20 times higher than the mechanical strength of the frame; and wherein the mechanical strength of the ceramic scaffold is the maximum load of the stress-strain curve.

43. The ceramic stent of any one of the preceding or following claims, wherein the coating polymer comprises a polymer formed by using n-butyl cyanoacrylate, 2-octyl cyanoacrylate, or a mixture thereof; wherein the mechanical strength of the ceramic support is in the range of 15N to 33N and/or the bending strength of the ceramic support is in the range of 10MPa to 50MPa when measured as the maximum load of a stress-strain curve; and wherein the bending strength is measured by a standard three-point bending test.

44. The ceramic stent of any one of the preceding or following claims, wherein the coating polymer comprises a polymer formed by using n-butyl cyanoacrylate, 2-octyl cyanoacrylate, or a mixture thereof; wherein the mechanical strength of the ceramic scaffold is at least 5 times greater than the mechanical strength of the frame, or at least 10 times greater than the mechanical strength of the frame, or 10 to 20 times greater than the mechanical strength of the frame; and wherein the mechanical strength of the ceramic scaffold is the maximum load of the stress-strain curve.

45. A ceramic stent according to any one of the preceding or following claims, wherein the coating substantially covers the at least one surface of the frame.

46. A method of manufacturing a ceramic stent, comprising:

preparing a green body using a three-dimensional (3D) printing method and the slurry;

debinding the green body and sintering to remove polymers formed by polymerization of the UV polymerizable monomer formulation to produce a sintered porous body, wherein the sintered porous body forms a framework;

coating the sintered porous body with a polymer coating solution;

Obtaining the ceramic support.

47. The method according to claim 46, wherein the three-dimensional (3D) printing method is a paste printing based on mask image projection (MIP-SP) method.

48. The method of any one of claims 46-47, wherein the green body is sintered at a temperature in a range of 1000 degrees Celsius to 1500 degrees Celsius.

49. The method of any one of claims 46-48, wherein the green body is sintered at a temperature in a range of 1050 degrees Celsius to 1250 degrees Celsius.

50. The method of any one of claims 46-49, wherein the coating of the sintered porous body is performed by a method comprising a surface spray coating method, a brush spray coating method, a vacuum fusion method, or a combination thereof.

51. The method of any one of claims 46-50, wherein the obtained ceramic scaffold is the ceramic scaffold of any one of claims 1-45.

52. The ceramic stent of any one of claims 46-51, wherein the coating polymer comprises a polymer formed by using surgical glue, a colloidal protein mixture, poly (ethylene glycol) dimethacrylate (PEGDMA), methacrylated gelatin (GelMA), gelatin, or a mixture thereof.

53. The method of any one of claims 46-52, wherein the coating polymer comprises a polymer formed by using a surgical glue.

54. The method of any one of claims 46-53, wherein the coating polymer comprises a polymer formed by using gelatin.

55. The method of any one of claims 46-54, wherein the coating polymer comprises a polymer formed by using cyanoacrylate glue, fibrin sealant, collagen-based compounds, glutaraldehyde glue, hydrogels, or mixtures thereof.

56. The ceramic stent of any one of claims 46-55, wherein the coating polymer comprises a polymer formed by using n-butyl cyanoacrylate, 2-octyl cyanoacrylate, or a mixture thereof.

57. The method according to any of the preceding or following claims, wherein the three-dimensional (3D) printing method is a paste printing based on mask image projection (MIP-SP) method.

58. The method of any one of the preceding or following claims, wherein the green body is sintered at a temperature in a range of 1000 degrees celsius to 1500 degrees celsius.

59. The method of any one of the preceding or following claims, wherein the green body is sintered at a temperature in a range of 1050 degrees Celsius to 1250 degrees Celsius.

60. The method of any one of the preceding or following claims, wherein the coating of the sintered porous body is performed by a method comprising a surface spray coating method, a brush spray coating method, a vacuum fusion method, or a combination thereof.

61. The method according to any one of the preceding or following claims, wherein the ceramic scaffold obtained is a ceramic scaffold according to any one of claims 1 to 45.

62. The method of any preceding or following claim, wherein the coating polymer comprises a polymer formed by using surgical glue, a colloidal protein mixture, poly (ethylene glycol) dimethacrylate (PEGDMA), methacrylated gelatin (GelMA), gelatin, or a mixture thereof.

63. The method of any preceding or following claim, wherein the coating polymer comprises a polymer formed by using a surgical glue.

64. The method of any one of the preceding or following claims, wherein the coating polymer comprises a polymer formed by using gelatin.

65. The method of any preceding or subsequent claim, wherein the coating polymer comprises a polymer formed by using cyanoacrylate glue, fibrin sealant, collagen-based compounds, glutaraldehyde glue, hydrogels, or mixtures thereof.

66. The method of any preceding or subsequent claim, wherein the coating polymer comprises a polymer formed by using n-butyl cyanoacrylate, 2-octyl cyanoacrylate, or a mixture thereof.

67. The method of any one of claims 46-66, wherein the degreasing is performed under vacuum at a pressure of 0.01-0.5 MPa.

68. The method of any one of claims 46-67, wherein the coating is applied to the stent under vacuum.

69. The method of any one of claims 46-68, wherein the coating is applied to the stent by surface spraying.

70. A method of replacing a bone of a subject, comprising:

71. A method of replacing a bone of a subject, comprising:

Placing the ceramic scaffold of any one of claims 1-45 at the location of the missing or removed bone portion, wherein the ceramic scaffold allows bone regeneration within and around the ceramic scaffold.

72. Any combination of the above claims is within the scope of what we claim.